[Show abstract][Hide abstract]ABSTRACT:
Photodynamic therapy (PDT) has recently emerged as a potential treatment alternative for head and neck cancer. There is
strong evidence that imprecise PDT dosimetry results in variations in clinical responses. Quantitative tools are likely to
play an essential role in bringing PDT to a full realization of its potential benefits. They can provide standardization of
site-specific individualized protocols that are used to monitor both light and photosensitizer (HPPH) dose, as well as the
tissue response for individual patients. To accomplish this, we used a custom instrument and a hand-held probe that
allowed quantification of blood flow, blood volume, blood oxygen saturation and drug concentration.

[Show abstract][Hide abstract]ABSTRACT:
Macroscopic modeling of singlet oxygen (1O2) is of particular interest because it is the major cytotoxic agent causing
biological effects for type II photosensitizers during PDT. We have developed a macroscopic model to calculate reacted
singlet oxygen concentration ([1O2]rx for PDT. An in-vivo RIF tumor mouse model is used to correlate the necrosis
depth to the calculation based on explicit PDT dosimetry of light fluence distribution, tissue optical properties, and
photosensitizer concentrations. Inputs to the model include 4 photosensitizer specific photochemical parameters along
with the apparent singlet oxygen threshold concentration. Photosensitizer specific model parameters are determined for
several type II photosensitizers (Photofrin, BPD, and HPPH). The singlet oxygen threshold concentration is
approximately 0.41 - 0.56 mM for all three photosensitizers studied, assuming that the fraction of singlet oxygen
generated that interacts with the cell is (f = 1). In comparison, value derived from other in-vivo mice studies is 0.4 mM
for mTHPC. However, the singlet oxygen threshold doses were reported to be 7.9 and 12.1 mM for a multicell in-vitro
EMT6/Ro spheroid model for mTHPC and Photofrin PDT, respectively. The sensitivity of threshold singlet oxygen
dose for our experiment is examined. The possible influence of vascular vs. apoptotic cell killing mechanism on the
singlet oxygen threshold dose is discussed using the BPD with different drug-light intervals 3 hrs vs. 15 min. The
observed discrepancies between different experiments warrant further investigation to explain the cause of the
difference.

[Show abstract][Hide abstract]ABSTRACT:
Selection and design of individualized treatments remains a key goal in cancer therapeutics; prediction of response and tumor recurrence following a given therapy provides a basis for subsequent personalized treatment design. We demonstrate an approach towards this goal with the example of photodynamic therapy (PDT) as the treatment modality and photoacoustic imaging (PAI) as a non-invasive, response and disease recurrence monitor in a murine model of glioblastoma (GBM). PDT is a photochemistry-based, clinically-used technique that consumes oxygen to generate cytotoxic species, thus causing changes in blood oxygen saturation (StO2). We hypothesize that this change in StO2 can be a surrogate marker for predicting treatment efficacy and tumor recurrence. PAI is a technique that can provide a 3D atlas of tumor StO2 by measuring oxygenated and deoxygenated hemoglobin. We demonstrate that tumors responding to PDT undergo approximately 85% change in StO2 by 24-hrs post-therapy while there is no significant change in StO2 values in the non-responding group. Furthermore, the 3D tumor StO2 maps predicted whether a tumor was likely to regrow at a later time point post-therapy. Information on the likelihood of tumor regrowth that normally would have been available only upon actual regrowth (10-30 days post treatment) in a xenograft tumor model, was available within 24-hrs of treatment using PAI, thus making early intervention a possibility. Given the advances and push towards availability of PAI in the clinical settings, the results of this study encourage applicability of PAI as an important step to guide and monitor therapies (e.g. PDT, radiation, anti-angiogenic) involving a change in StO2.

Data provided are for informational purposes only. Although carefully collected, accuracy cannot be guaranteed.
The impact factor represents a rough estimation of the journal's impact factor and does not reflect the actual
current impact factor.
Publisher conditions are provided by RoMEO. Differing provisions from the publisher's actual policy or licence
agreement may be applicable.

probe, yielding useful data only between light fractions. Opticalcoherence tomography has high spatial and temporal resolu-tion, permitting the imaging of local vascular changes (18, 19).However, optical coherence tomography can only measuresmall near-surface tissue volumes (f1 mm2; ref. 20), similar tothe laser Doppler technique. Power Doppler ultrasound cannoninvasively follow changes in tumor perfusion after PDTthrough determination of the color-weighed fractional area inacquired images (21). However, it does not readily allowcontinuous measurement during PDT.Near-IR diffuse correlation spectroscopy (DCS) enablesmeasurement of rBF noninvasively through deep tissues. Ithas been successfully applied and validated in studies offunctional imaging and spectroscopy of brain (22–24), tumorphysiology (25), tissue burns (26), and exercise medicine(27, 28). DCS is closely related to the commonly used laserDoppler technology, but it offers the advantage of deepermeasurement through tissues. Furthermore, with appropriateoptical cutoff filters, DCS can be used to follow tumor bloodflow continuously during PDT.In this study, we employed a DCS system and a uniquenoncontact probe to monitor the rBF of murine tumors duringillumination for Photofrin-PDT and at specific time points aftertreatment. Within minutes of beginning PDT, rBF rapidlyincreased, followed by a decline and subsequent peaks anddeclines of various kinetics. The slope (flow reduction rate) andduration (interval time) over which rBF decreased followingthe initial PDT-induced increase was highly associated withtreatment durability; treatment durability is measured as thetime of tumor growth to a volume of 400 mm3(time-to-400mm3). After PDT, all animals showed decreases in rBF at 3 and6.5 hours, and rBF at these time points was also predictive oftumor response. Broadband diffuse reflectance spectroscopywas employed to measure the tissue hemoglobin oxygensaturation (SO2) of the PDT-treated tumors, finding that SO2decreased after PDT in association with the decreases in rBF.These data show that DCS-measured changes in tumor rBFduring and after Photofrin-PDT are predictive of treatmentefficacy.Materials and MethodsTumor model and photodynamic therapy.sarcoma tumors were propagated on the shoulders of C3H mice(Taconic, Germantown, NY) by the intradermal injection of 3 ? 105cells. Animals were treated f1 week later when tumors were f6 to 7mm in diameter. The photosensitizer Photofrin (Axcan Pharma, Inc.,Mont-Saint-Hilaire, Quebec) was given via tail vein at 5 mg/kg at f24hours before illumination. The laser system consisted of a KTP YAGpumped dye module (Laserscope, San Jose, CA) tuned to produce 630nm light. Light was delivered to a 1-cm-diameter treatment fieldthrough microlens-tipped fibers (CardioFocus, Norton, MA) and thepower density was measured with a power meter (Coherent, Auburn,CA). Treatment was to a total fluence of 135 J/cm2, delivered at75 mW/cm2. Treatment groups consisted of animals receiving PDT(n = 15) and animals receiving light but no Photofrin (n = 10). Toalter blood flow in a manner independent of PDT, an additional11 animals received an i.v. injection (2.5 or 5.0 mg/kg in saline) ofhydralazine (Sigma Aldrich, St. Louis, MO) at 10 minutes beforebeginning illumination. During illumination and during opticalmeasurement, mice were anesthetized with isoflurane and kept warmon a heating pad. Treatment durability was measured as the number ofdays after PDT or after control treatment until tumor growth to aRadiation-induced fibro-volume of 400 mm3(tumor volume = length ? width2? 3.14/6; i.e.,time-to-400 mm3).Diffuse correlation spectroscopy.flow continuously during PDT (from 10 minutes before PDT until15 minutes after PDT) and for 10-minute durations at times 3 and6.5 hours after PDT. In animals that received hydralazine, blood flowmonitoring began 10 minutes before hydralazine injection andcontinued through another 10 minutes preillumination, 30 minutesof illumination, and 15 minutes after PDT. A 3-hour time point wasalso evaluated in mice treated with hydralazine and PDT. The majorcomponents (25) of the DCS system are the following: an 800-nm lasersource with long coherence length, operating in continuous mode; anoncontact probe with 13 source and four detector fibers; and a camerato deliver excitation light and collect reflected diffuse light, respectively(Fig. 1A), from the tissue surface. The noncontact probe (Fig. 1B)consists of 13 source fibers arranged in concentric circles within a 6-mmdiameter. Using optical switches, the 800-nm light was directedthrough each of the source fibers in a consecutive manner; thesampling time for one scanning frame (i.e., all source-detector pairs)was 18 seconds. Four single photon–counting avalanche photodiodeswere employed in parallel for detection of the diffuse light. The probehead was mounted behind a camera lens fixed at a distance of 15 cmfrom the tumor. This setup enabled us to monitor blood flow duringPDT by permitting unobstructed illumination with the treatment lightat a small angle to the tissue surface. A long-pass optical filter (03FCG507, Melles Griot, Rochester, NY) in front of the camera lens attenuatedlight below 650 nm, preventing the 630-nm treatment light fromsaturating the detectors.DCS was used to monitor bloodwww.aacrjournals.org Clin Cancer Res 2005;11(9) May1, 2005 3544Fig.1. A schematic of the DCSinstrument (A) anda map (B) of the probe formeasuringtumorbloodflow.Opticalfibersfor13sourcepositions(smallcircles)andfour DCS detectors (squares) are arrangedina two-dimensionalpattern. Largedashed circle, contourof the tumor.Cancer Therapy: Preclinical

Page 3

Speckle fluctuations of the diffuse light are sensitive to the motionsof tissue scatterers such as red blood cells. The quantity containing thisinformation is the electric field (E(r,t)) (26). The electric field temporalautocorrelation function, G1(r,s) = hE(r,t) E*(r,t + s)i, or its FourierTransform can be explicitly related to the motion of the scatterers (e.g.,red blood cells) even in turbid media (29–31). Here the angle bracketsh i denote averages over time and s is called the correlation delay time.A continuous wave laser with a long coherence length and a single-photon counting avalanche photodiode are needed for DCS measure-ments. An autocorrelator takes the avalanche photodiode output anduses photon arrival times to compute the light intensity temporalautocorrelation function. From the normalized intensity autocorrela-tion function, we calculate the normalized electric field correlationfunction, g1(r,s) = G1(r,s)/hE(r,t) E*(r,t)i; G1(r,s), satisfies thecorrelation diffusion equation in highly scattering media (26, 31).The exact form of the correlation diffusion equation depends on thenature and heterogeneity of the particle motion. For the importantcase of random flow in the tissue vasculature, the mean-squaredisplacement, (hDr2(s)i), of the scattering particles (e.g., blood cells)in time s is hDr(s)i = hV2is2. Here hV2i is the second moment of thecell velocity distribution. For the case of diffusive motion, hDr2(s)i =6DBs, where DBis an effective diffusion coefficient of the movingscatterers. We have found that both of these models fit our data, butthe latter model often provides better quality fits (23). In this case thenormalized correlation function g1(r,s) will decay approximatelyexponentially in s. Its decay depends on a variable a (proportionalto the tissue blood volume fraction) and on the mean-squaredisplacement of the blood cells. Relative changes in DB orare correlated with relative changes in blood flow. In principle, DCScan measure absolute levels of blood flow. In practice, however, itis desirable to calibrate DCS with other measurement techniques(e.g., Arterial Spin Label magnetic resonance imaging, ultrasoundDoppler, etc.) for quantification of absolute flow. In this study, onlyrelative blood flow is reported, and these relative blood flow changesare insensitive to our model for hDr2(s)i (i.e., the relative change ofDBandffiffiffiffiffiffiffiffiffiffihV2iconcepts and approximations can be found in the noted references(22, 23), and a comparison with Power Doppler ultrasound will beprovided in this article.DCS reflects local flow in primarily small vessels (i.e., arterioles,capillaries, and venules). From diffusion theory, the maximumpenetration depth of diffuse light in tissue depends on tissue opticalproperties and source-detector separation. The most probable penetra-tion depth of diffuse light in tissue is roughly one third to one half ofthe source-detector separation on the tissue surface. Therefore, a specificsource-detector pair predominately provides information about aparticular tissue layer. The source-detector separations used in thisstudy ranged from 1 to 3.5 mm, providing flow information fromdepths of 0.5 to 1.7 mm. We found the magnitude of PDT effect on rBFdiffered with tumor depth, although the trends were similar amongtumor layers. In this study, DCS data for one scanning frame wereaveraged over all source-detector separations (1-3.5 mm), representingthe rBF in the bulk tumor tissue. rBF at 3 and 6.5 hours were obtainedby averaging over 10 minutes of monitoring. rBF during PDT wascharacterized by the variables of a linear regression model fit to the firsttreatment-induced decrease in flow. The SE of the slope and intercept ofthe linear model represent the fitting errors in flow reduction rate andinterval time, respectively. To assess instrument reproducibility,Brownian motion of particles was measured in an Intralipid (LiposynIII, Abbott Laboratories, Chicago, IL) phantom (1%) with knownreduced scattering coefficient (lsV = 10 cm?1) and absorptioncoefficient (la= 0.02 cm?1; ref. 32) at 785 nm. The coefficient ofvariation of multiple measurements (n = 10) with multiple separations(1-3.5 mm) was <3%.Broadband diffuse reflectance spectroscopy.properties and determine tissue hemoglobin oxygen saturation (SO2),broadband reflectance spectrometric measurements were made at timesffiffiffiffiffiffiffiffiffiffihV2ippare almost identical). A detailed description of theseTo quantify tissue optical15 minutes before PDT and 20 minutes, 3, and 6.5 hours after PDT. Thebroadband reflectance spectrometer collected reflectance spectra,determined by tissue scattering and absorption, at many source-detectorseparations. Based on the absorption data, tissue concentrations ofchromophores such as oxyhemoglobin (HbO2) and deoxyhemoglobin(Hb) were calculated. This instrument has been described in detail inrecent publications (33, 34). Briefly, the system consists of a 250-Wquartz tungsten halogen lamp (Cuda Fiberoptics, Jacksonville, FL), ahand-held surface contact fiber optic probe, a spectrograph (SpectraPro-150, Acton Research, Acton, MA), and a liquid nitrogen–cooled CCDcamera (LN/CCD-1100-PF/UV, Roper Scientific, Trenton, NJ) to imagethe reflectance spectra from multiple detection fibers simultaneously.The fiber optic probe consisted of a 400-Am-diameter source fiber and10 colinear 400-Am-diameter detection fibers at various source-detectorseparation distances. Due to poor signal at larger source-detectorseparation distances and failure of the diffusion model at very smallsource-detector separation distances, only signals from the detectionfibers with the second and third shortest source-detector separationdistances, 1.2 and 1.8 mm, were used in our algorithm to calculatetumor optical and physiologic properties, including HbO2and Hb.From this information, the total hemoglobin concentration (THC =cHbO2+ cHb) and the tissue hemoglobin oxygen saturation (SO2 =cHbO2/THC) were calculated. SO2typically had a coefficient of variationof <1.5% in a single location and f2% to 7% when averaged over thewhole tumor. The latter is a function of both instrument-introducedvariability (<1.5%) and heterogeneous oxygen distribution within themeasured tissue (tumor). In previous studies, the broadband reflectancespectrometer was validated through the creation of the Hill curve for atissue phantom of human or mouse erythrocytes over the course of itsdeoxygenation. The difference between the smoothed fit to themeasured oxygen dissociation curve and published values was <5%(33, 34).Power Doppler Ultrasound Imaging.imaging of tumor perfusion was done with a broadband 12-5 MHztransducer using a Philips ATL 5000 (Philips ATL, Bothell, WA)ultrasound scanner. Imaging was done under ketamine/xylazineanesthesia (150/10 mg/kg) at times 15 minutes before, 15 minutesafter, 3 hours after, and 6.5 hoursafter PDT; mice were kept warm ona heating pad. Over a period of 5 to 10 minutes, f10 images wereacquired of each tumor. The pulse repetition frequency, color gain, andwall filters were held constant for all images. In power Doppler images,tissue regions with blood flow are coded in color. The hue, brightness,andvalueofthecolorrepresentthestrengthoftheDopplersignalandarerelated to the concentration of the moving RBC. The color level isexpressed in arbitrary units from 0 to 100, where the values 0 and 100representnopowerDopplersignalandmaximumpowerDopplersignal,respectively.Oneachimage,thetumorwasidentifiedandoutlinedastheregion of interest based on a grayscale scan. Within the region of interest,the color-weighed fractional area was calculated as the product ofthe fractional area (ratio of colored to total pixels in the image) and themean color level (the sum of the integrated power values divided bythe number of colored pixels) as described earlier (35, 36). The averagecolor-weighed fractional area was calculated from all images for eachmouse at each time point and studied as a measure of tumor perfusion.Because mean color level represents RBC flux and fractional area coveredby the colored pixels represents the area of perfusion, the productof the two is proportional to the blood volume moving through theimage plane.Statistics.Statistical analyses were carried out using R v.1.70 (freesoftware, http://www.r-project.org) and Microsoft Excel 2000. Fordescriptive purposes, figures display means and SEs or SDs based onthe individual data collected at each time point. However, for thepurposes of statistical testing, we used SEs based on data modeled overall time points using the methods described below. Mixed effectsmodels were used to compare blood flow patterns in animals measuredby DCS versus power Doppler (37). The mixed effects model isan analysis of variance in the sense that it accounts for multiplePower Doppler ultrasoundwww.aacrjournals.orgClin Cancer Res 2005;11(9) May1, 20053545Noninvasive Blood Flow Monitoring During PDT

Page 4

experimental factors as well as allowing for correlations betweenrepeated measurements on the same animal. For the comparison ofDCS and power Doppler, the model included time, modality (powerDoppler versus DCS), and treatment (control versus PDT) as maineffects and included interaction terms between time and treatment toallow the effect of PDT to differ over time. Based on the AkaikeInformation Criteria, separate random effects were used to model theinteranimal variance for animals measured with power Doppler andDCS, with a common random error term to model the intra-animalvariance (37). Separate analyses were also carried out to assess theeffects of time and treatment on SO2and rBF. Likelihood ratio testswere used to assess the significance of the differences in SEs amongpower Doppler and DCS-measured animals and to assess thesignificance of the grouping factors (e.g., DCS versus power Dopplerand PDT versus control; ref. 38). We used the results from the model toestimate average outcome for each time point for control and PDT-treated animals and as the basis of a Wald test comparing variables(rBF and SO2) at individual time points to their baseline. All tests weretwo sided and used a type I error rate of 0.05.The association between rBF and SO2 or time-to-400 mm3wasassessed for the best-fit model (linear, log, or exponential), as de-termined by the correlation coefficient (r2). The statistical significanceof this association was determined by a Wald test comparing the slopeof the model to a slope of zero.The Wilcoxon rank sum test was used to compare rBF or treatmentdurability between animals with and without hydralazine.ResultsDiffuse correlation spectroscopy and power Doppler ultrasoundmeasure similar changes in relative blood flow after photodynamictherapy. DCS measurement of tumor blood flow was validatedthrough comparison with power Doppler measurement ofmoving blood volume in tumors treated with the same PDTtreatment conditions (5 mg/kg Photofrin, 135 J/cm2, 75 mW/cm2). Data from both instruments were expressed as the rBF(i.e., as a percentage of the baseline value measured over the 10minutes before PDT). Figure 2 shows that DCS and powerDoppler ultrasound detected similar changes in tumor bloodflow at times 15 minutes, 3, and 6.5 hours after control (Fig. 2A)or PDT (Fig. 2B) treatment. The model-based estimate of themean difference between rBF measured by DCS versus powerDoppler was 11.5% (95% confidence interval, ?19.0 to 42.0),a difference which was not statistically significant (P = 0.45)and which suggested overall good agreement between theaverages estimated by the two modalities.Tumor blood flow fluctuates significantly during photodynamictherapy. Compared with power Doppler, DCS offers the ad-vantage of continuous blood flow monitoring during thedelivery of the 630-nm illumination for Photofrin-PDT. UsingDCS, relative blood flow was measured continuously beginning10 minutes before the start of PDT until 15 minutes after theconclusion of PDT. Rapid changes in rBF were detected over thecourse of the 30-minute PDT treatment. Averaged traces of rBFduring control or PDT treatment of animals are shown in Fig. 3.In unphotosensitized controls (Fig. 3A), minor fluctuations inrBF were detected during illumination. In PDT-treated animals(Fig. 3B), a rapid increase in rBF occurred during the first10 minutes of treatment, peaking at 168.1 F 39.5% (average FSE) of the baseline. This increase was remarkably consistentamong animals, and the peak occurred within the first f1 to10 minutes of treatment (range, 0.9-10.5 minutes). Followingits PDT-induced increase, rBF decreased to 59.2 F 29.1%(average F SE) of baseline. On average, the first minimum wasreached at 16 minutes after the start of treatment, but the timeto the trough varied widely among tumors (range, 4.0-24.2minutes). Tumors that reached this first trough at shorter timesafter the beginning of treatment tended to exhibit subsequentdistinctive peaks and declines in rBF. However, these additionalpeaks were of a smaller magnitude than the initial increase inrBF. Representative examples of a tumor that showed a singlepeak in rBF versus one that showed three peaks are shown inFig. 3C and D, respectively.To characterize the PDT-induced changes in rBF duringillumination, the following variables were defined: the maxi-mum differential flow (rBFmax? rBFmin), interval time (Tmin?Tmax), and flow reduction rate (maximum differential flow /interval time). Here the rBFmaxand rBFminare the maximum andminimum flow of the first peak, respectively (see Fig. 3C and D).Tminand Tmaxare the time points when flow reaches the rBFmaxand rBFmin. Analysis of animals with multiple peaks in rBF wasbased on the first peak because it was the most consistent, aswell as generally the highest. In PDT-treated animals, maximumdifferential flow ranged from 61.8% to 137.1%, interval timeranged from 2.2 to 15.6 minutes, and flow reduction rate rangedfrom 4.4 to 45.8 minute?1. Thus, substantial variability waswww.aacrjournals.org Clin Cancer Res 2005;11(9) May1, 2005 3546Fig. 2. rBF measured by DCS (o) and power Doppler ultrasound (4) at 15minutes, 3, and 6.5 hours after control (A) or PDT (B) treatment. Photofrin-PDTdone to135 J/cm2at 75 mW/cm2; controls receivedillumination, but no Photofrin.rBFat each time point was calculated as the percentage of the baseline value,measuredinthe same animalover the15 minutes before PDT. Points, average;bars, FSE. Fifteentreated and10 controlanimals were evaluatedby DCS; fivetreatedand 5 controlanimals were evaluatedby power Doppler (at 3 hours, onlyfouranimals were available).Cancer Therapy: Preclinical